Imaging technology using an optical device has been prevailing not only in commercial electronic devices such as cameras, printers and facsimile machines but also in the medical field. X-ray photography using X-rays and diagnosis using ultrasonic waves have been widely used so far in order to non-invasively image slices in vivo. Because of exposure problems, a method using X-rays has remarkable restrictions on usage frequency and biological parts to which the method is used, and further, the resolution thereof is restricted to that of film shooting at the same magnification. A method using ultrasonic waves has no exposure problem, and therefore, has no such usage restriction as with X-rays. However, the resolution thereof is normally nothing more than roughly 1 cm. Therefore, it is impossible for the method to perform imaging at a cellular-level size.
In medical settings, a new technology has been demanded whereby images of slices under the epidermis of a living organism can be generated at a resolution of micron order. Optical coherence tomography (hereinafter referred to as OCT), since it was developed in the 1990s, has been known as a technology for implementing this.
OCT utilizes the principle of a Michelson interferometer. Low coherence light used was irradiated to a living organism. Images under the epidermis of a living organism are obtained based on interfering light produced by reference light and reflected light from the living organism. OCT has been practically used as a diagnosis device indispensable for ophthalmology in a retinal diagnosis.
FIG. 13 is a diagram explaining the basic principle of OCT (see NPL1). Only the outline of the basic principle will be hereinafter briefly described. Low coherence light with coherence length Δlc is supplied as incident light to a living organism 4 from a low coherence light source 1. Outgoing light 6 from the light source 1 enters a beam splitter 2 and is therein split into two halves. Light 7, which is one of the two split halves of light, travels to a movable mirror 3. The light 7 is reflected by the movable mirror 3, and travels as reference light 8 again towards the beam splitter 2. Light 9, which is the other of the two split halves of light, is reflected by any of reflection surfaces A, B and C located at different depths inside the living organism 4, and signal lights 11a, 11b and 11c are respectively obtained. Each signal light interferes with reference light 10 via the beam splitter 2. Due to the interference, such reflected light as having a wavefront distorted by excessive scattering within the living organism is removed, while such reflected light as maintaining an original plane wave is only detected selectively.
Here, when the movable mirror 3 is located in a position A′, interference is only caused by the light reflected by the reflection surface A. At this time, the reference light and the signal light interfere with each other and an electric signal is obtained from a photo detector 5, when the relation of the following equation is satisfied where the distance between the center of the beam splitter 2 and the movable mirror 3 is set as LR while the distance between the center of the beam splitter 2 and the reflection surface A is set as LS.|LR−LS|<Δlc  equation (1)
The above equation is established on a one-to-one basis respectively: between the reflection surface A and the position A′ of the movable mirror; between the reflection surface B and a position B′ of the movable mirror; and between the reflection surface C and a position C′ of the movable mirror. Therefore, it is possible to measure a reflected light intensity distribution under the epidermis of the living organism along an optical axis (z-axis) direction within the living organism 4 with a spatial resolution Δlc by continuously moving the movable mirror 3 at a constant speed v. It is possible to obtain a reflected light intensity distribution under the epidermis of the living organism within an x-z plane by scanning incident light into the living organism along an x-direction by means of a scanning mirror or the like, and a final OCT image is generated based on the obtained distribution. An optical fiber coupler is usable for the configuration of the Michelson interferometer in FIG. 13, and thus, a testing device sufficiently usable even in medical settings is implemented.
In OCT configured as shown in FIG. 13, image data is obtained on a time-series basis by moving the movable mirror. Therefore, this type of OCT is referred to as time domain (TD)-OCT (hereinafter referred to as TD-OCT). In TD-OCT, it is required to move a movable mirror having a mass, and therefore, the scan speed has a limitation. However, it is often difficult to completely fix a living organism in some testing conditions, and it is preferable to obtain necessary image data by performing a scan in a short time as much as possible. Incidentally, when a scan is performed for a part in vivo that is located at a depth where blood vessels exist, it becomes difficult to obtain images because of scattering by red blood cells moving within the blood vessels. Therefore, it has been demanded in OCT to obtain information in the depth direction as quick as possible.
In view of the above, Fourier domain (FD)-OCT (hereinafter referred to as FD-OCT) was proposed for obtaining reflected light intensity along an optical axis by Fourier-transforming an interfering signal. In FD-OCT, a spectrometer, decomposing a signal light from a living organism into lights with respective wavelengths, is disposed forwards of the photo detector 5 in FIG. 13, and a detector factor array of the photo detector detects an interfering spectrum. In other words, the interfering spectrum is obtained by a parallel detector in which a large number of detector factor elements relevant to the respective wavelengths are disposed. A reflected light intensity distribution along an optical axis is obtained by Fourier-transforming the spectrum detected by the parallel detector. However, it is required in FD-OCT to provide the parallel detector having a large number of detector factor elements. When a large number of detector factor elements are provided, it is required to simultaneously detect signals with respective wavelengths in more than 1000 detector factor elements. Such parallel detector has been implemented at a wavelength band of 1.1 μm or less in the form of a silicon CCD or a CMOS, but it is difficult to obtain a parallel detector implemented at a longer wavelength.
Therefore, at this point of time, FD-OCT is applicable to a retinal diagnosis for which visible light is usable, but is not applicable to OCT for tissues (e.g., skin) requiring an operation at a longer wavelength range. Further, in application to slice imaging of a blood vessel, absorption by hemoglobin of red blood cells cannot be ignored unless light with a longer wavelength of roughly up to 1.3 μm is used because of scattering by hemoglobin. On the other hand, absorption by water becomes remarkable, in turn, when the wavelength of a light source gets closer to 1.5 μm. It becomes difficult to obtain a photo detector when the wavelength exceeds 1.6 μm. Due to these reasons, it has been demanded to use a light source at 1.3 μm band in order to utilize OCT for skin and so forth.
In view of this, a method having received attention anew is swept source (SS)-OCT (hereinafter referred to as SS-OCT) for sweeping the frequency of a light source by changing FD-OCT. In SS-OCT, a light source wavelength is regularly swept unlike FD-OCT in which a large number of wavelength signals are generated at one time by causing a spectrometer to decompose signal light obtained by irradiating a living organism with light from a coherence light source. Through the sweep of the frequency of light from the light source, signals with respective wavelengths can be detected on a time-division manner using a single detector. In other words, in FD-OCT, wavelength division is performed based on a spatial position by a spectrometer, whereas in SS-OCT, wavelength division is performed based on time and therefore only a single detector is required. A parallel detector having a large number of detector factor elements is not required, and there is no limitation on selection of a detector. Therefore, even a light source at a 1.3 μm band can be used.
FIG. 14 is a diagram schematically shown for explaining the principle of SS-OCT (NPL1). In SS-OCT, an optical signal 26 that the optical frequency thereof is linearly swept with respect to time is supplied to a living organism 24 from an optical frequency (wavelength) swept light source 21. For example, a wavelength variable laser is used as the optical frequency swept light source 21. In SS-OCT shown in FIG. 14, the position of a mirror 23 is stationary. Respective elements are disposed for establishing LR=LS, where the distance between the center of a beam splitter 22 and the mirror 23 is set as LR while the distance between the center of the beam splitter 22 and a living organism surface 31 is set as LS.
At this time, regardless of time, differences in optical frequencies are constant between reference light 28 and reflected lights 29b and 29c respectively from reflection surfaces 32 and 33 in vivo. Where these differences in optical frequencies are set as f2 and f3, signal light, in which beat frequencies f2 and f3 relevant to the reflection surfaces 32 and 33 coexist, is obtained by the interference between the reference light 28 and the reflected lights 29b and 29c. When the signal light is Fourier-transformed, reflected light intensities at the beat frequencies f2 and f3 are obtained. When the optical frequency from the light source 21 is linearly swept, the beat frequencies f2 and f3 are directly proportional to depths d2 and d3, respectively. Reflected lights are produced at respective positions in vivo, and therefore, it is possible to obtain a distribution of reflected light intensity along the optical axis (z-axis) direction by Fourier-transforming the interfering light. When a beam scan is also performed in the x-axis direction, an OCT image within an x-z plane is obtained.
Unlike the well-known FD-OCT, no parallel detector is required for SS-OCT because only a photo detector 25 is required to detect signal light in which interfering lights at different beat frequencies coexist by means of a single detection element. It becomes possible to use a swept light source at a 1.3 μm band that is preferred to a diagnosis of skin and so forth. Practical use of SS-OCT has been progressed even in the fields other than ophthalmic care due to: its stable configuration using a photo fiber coupler; its high-speed image obtainment attributed to inessential of a movable mirror; and its easiness in utilizing various photo detectors.